According to the prior art, positron emission tomography detector rings are used in order to detect the ß+ß− annihilation radiation. The rings consist of scintillation crystals which are adjoined by photosensors which are capable of detecting the scintillation radiation. Typical photosensors are photomultipliers (PMT), avalanche photodiodes (ADP), photodiodes, and silicon photomultipliers (SiPM). The design is such that the detector ring is generally circular, wherein the object to be measured, e.g., a body part of a patient or animal, is placed in the center of the detector ring (PET ring). The use of radiodiagnostic agents generates ß+ß− annihilation radiation which is to be detected. The ß+ß− annihilation radiation impinges on scintillation crystals which are arranged annularly or squarely around the object to be examined and generate the scintillation radiation. The scintillation radiation is in turn registered by the photosensors which, in relation to the radiation source, are located in the concentric arrangement behind the scintillation crystal. However, the photosensors can also be arranged on other sides of the scintillation crystal—for example, in front of the scintillation crystal or laterally thereto. The scintillation crystal is a three-dimensional body. Based upon an arrangement in which the object to be examined emits annihilation radiation from the center of the detector ring, the cross-section on which the annihilation radiation impinges on the scintillation crystal spans an x-y axis. The depth of the scintillation crystal is referred to in this nomenclature as the z-axis. In an idealized representation, an object to be examined or an emission source for radiating an energy of 511 keV is located in the center of the detector ring, which energy ideally impinges perpendicularly on the x-y plane of the scintillation crystal and has a penetration depth along the z-axis of the scintillation crystal. The 511 keV annihilation radiation then triggers a scintillation at a point of the scintillation crystal along the z-axis, which scintillation is registered as a signal by the photosensor—for example, an SiPM. An SiPM is capable of detecting even individual photons.
There is a correlation between the sensitivity of the scintillation crystal and its length along the z-axis. The deeper the scintillation crystal is dimensioned, the more sensitive it is, since a scintillation event is the more likely to occur. In the detection of the annihilation radiation, rays are emitted in two opposite directions from the point at which the annihilation radiation is emitted, so that the rays form an angle of 180°. The line formed by these rays is referred to as the “line of response” (LOR). Accordingly, in the case of an annular detector, two rays impinge along the LOR on scintillation crystals, which, based upon the annular arrangement in whose center the emission source is located, are on opposite sides.
For photodetectors with light detection on only one side of the scintillation crystal, various established methods exist for determining the x- and y-position of an event. However, these do not include the z-position, and the exact position in the scintillation crystal where the gamma photon was stopped on the z-axis and converted to light is thus not determined. If the z-position is not determined as well, parallax errors, which are attributable to the so-called depth-of-interaction problem (DOI problem), occur in the determination of the LOR. The DOI problem occurs whenever the point from which the emission of the annihilation radiation emanates, in an annular detector, is not exactly in the center. The farther the emission center for an LOR is outside the center of a PET ring, the greater the problem becomes. In the design of a PET ring, this leads to a compromise between increasing the sensitivity through longer scintillation crystals and reducing the DOI errors through shorter scintillation crystals. In some areas of the PET application, there is the need to use PET rings (detector rings) closely adjacent to the examination object. This is particularly the case in medicine when patients are to be examined simultaneously with an MRI method and a PET method. In these hybrid scanners, the PET ring must fit into the opening of the MRI scanner tube. The PET ring used must consequently be dimensioned to be small in diameter so that it fits into the opening of the MRI ring. However, with a small dimensioning of the PET ring, there is the problem that the object to be examined, e.g., a body part of a small animal or even a human, can be arranged to be centered, but is dimensioned, when measured at the diameter of the PET ring, such that it reaches far into the edge regions of the opening of the PET ring. However, points from which annihilation radiation emanates are thus also positioned so close to the PET ring that the DOI problem becomes significant.
In past years, the resolution in small animal PET scanners, in particular, was significantly improved by the use of pixelated scintillation crystal blocks with ever smaller pixel sizes. In this case, the pixilation is realized on the x-y plane, so that tubes of pixels which are oriented in the z-direction are formed in the scintillation crystal. This was prompted especially by the need for ever higher spatial resolution in small animal PET scanners, since the examined object is very small. In the meantime, the pixel size has already reached the sub-millimeter range. This has amplified two problems which need to be solved. First, the pixelated crystal blocks consist of adhesive and reflector film which are located between the individual scintillation crystals in order to thus construct the pixelated block. The layer of adhesive and reflector film has an approximate thickness of 70 μm. Accordingly, pixelated arrays with particularly low pixel spacing have an increased sensitivity loss. In the case of an array with crystal pixels of 0.8 cm×0.8 cm size, as were used, for example, in [1], the ratio of adhesive and film to scintillation crystal is significantly reduced, so that the adhesive and film already make up a proportion of 29%. The scintillation crystal proportion is consequently reduced to 71%. In the other 29% of the volume, gamma quanta cannot be stopped and converted to light. If even smaller pixelated arrays of, for example, 0.5 cm×0.5 cm are used, the crystal proportion is even reduced to 59%. Therefore, the increase in resolution with pixelated arrays is always associated with a loss of sensitivity. The second problem with pixelated scintillation crystal arrays is that the emitted light is concentrated on a smaller area of the photosensor surface. This is a problem particularly for binary photosensors, such as SiPM's. An SiPM consists of several microcells which function as binary elements. They detect whether light has been detected or not. When light is detected, the microcell performs a breakthrough. The number of broken through microcells quantitatively indicates how much light has reached the detector surface. When two or more light quanta trigger a microcell, the output signal remains the same. The more light that hits an SiPM, the higher the probability of two or more light quanta impinging on the same microcell of the SiPM. These additional light quanta cannot be detected. Consequently, the probability of saturating a microcell is significantly higher when pixelated scintillation crystal arrays are used, since these arrays concentrate light more strongly on a small area of the photosensor. Saturation effects also lead to poorer energy resolution of the detectors.
Prior art detectors use SiPM-based photosensor technologies in order to enable magnetic resonance tomography (MRI) compatibility for use in MR/PET hybrid scanners. Another problem with hybrid scanners is that the space for PET detectors and associated electronics is limited by the tube diameter of the magnetic resonance imaging scanner (MRI). This applies, in particular, to ultra-high-field scanners. As a consequence of the narrower tube diameters, the PET scintillation crystals must be as short as possible. Shorter scintillation crystals also reduce the sensitivity. This also means that the PET ring is located closer to the examination object due to the requirements of the tube diameter. The parallax error is larger the closer to the PET ring the annihilations, and thus the resulting LOR, take place. This is because the gamma quanta no longer pass perpendicularly into the scintillation crystals when the annihilation occurs close to the PET ring. In the PET ring design, this has the consequence that the parallax errors increase and become stronger when the PET ring is close to the object to be examined, since annihilation can likewise occur close to the PET ring in this case. Apart from limitations by hybrid devices, it is also attempted to design the PET rings as small as possible due to higher sensitivity and lower cost.
Furthermore, it is known that photosensor concepts can include coding of the output channels, since the power consumption of the PET ring is increased by increasing the output channels. This is, however, limited by the design. A simple calculation illustrates this. A PET ring with a diameter of 8 cm and a length of 10 cm results in a detector surface area of 251 cm2. If a 1-to-1 coupling of scintillation crystals and photosensors with a crystal pixel size of 0.8 mm is used, 39,270 readout channels are already needed if each channel is read out individually.
In order to achieve higher spatial resolutions, current sensor designs consist of sensor chips with narrower pixel sizes. This leads to a significant increase in the readout channels, which are limited by the power consumption, space, and data rates. As a consequence, position-sensitive (PS) coding methods were developed to reduce the number of readout channels of a photosensor [1-6].
A concept published in [7] proves the possibility of constructing a PET detector consisting of monolithic crystals and SiPM's. As already mentioned above, monolithic crystals solve the problem of sensitivity losses due to the space requirement of reflector films and associated adhesives. As a result, the production costs of monolithic crystals are lower. The thickness of the crystals used is 2 mm. The design used in [7] thereby reduces parallax errors, which is, however, paid for by the small extension of the scintillation crystal in the z-direction. At the same time, the detection efficiency is, however, low due to the low crystal height.
There are various possibilities for measuring DOI information and thus correcting parallax errors, which additionally detect light on a further crystal side. The costs are thereby immensely increased—particularly for prior art SiPM's. A concept for DOI detection which detects light only on one crystal side and uses monolithic crystals is published in [8] and patented in [9]. It uses the known principle that the light distribution of the crystal is dependent upon the DOI. The detector concept used is coupled with monolithic crystals to position-sensitive photomultiplier (PMT) H8500 from Hamamatsu. A resistor network is also used, which enables position coding and thus also output channel reduction. In this case, the standard deviation of the light distribution is used to estimate the DOI. In order to calculate the standard deviation, the moment of the 1st and 2nd orders of the light distribution is required. The moment of the 1st order is already given by the linear coding of the output channels. In order to determine the moment of the 2nd order, a sum network has been developed and integrated into the resistor network. This significantly increases the complexity of the sensor chip.
An overview of PET detectors with DOL detection is summarized in [10]. Descriptions and results of small animal PET and MR/PET hybrid scanners, which have been developed in recent years, are contained in [11-14].
The detector described in [7] is realized with monolithic crystals. A closely adjoining ring was designed, to increase sensitivity. Monolithic crystals were used at the same time. Due to the resulting short distance between the scintillation crystals and the examination object, the DOI problem is increased. The developers of the ring are therefore limited to a 2 mm crystal thickness. As a result, the sensitivity gained by the narrow ring and the use of monolithic crystals is lost again by the small thickness of the scintillation crystals. However, this work proves that high resolution is possible with monolithic crystals.
DOI positions can be determined by attaching sensors to two crystal faces. This requires double the photosensor surface. Currently, sensors are one of the most expensive components of a PET ring.
A three-dimensional animal PET scanner was integrated into a 7T animal scanner by Judenhofer et al. [11]. It is based upon APD's which use scintillation crystals having a thickness of 4.5 mm and consist of crystal arrays having 144 crystals and a distance of 1.6 mm. The crystal array is coupled to a 3-by-3 APD array. The axial field of view (FOV) is 19 mm. This developed system shows that space is greatly limited especially for integrated systems, which forces a compromise between crystal thickness and axial FOV. This results in the low sensitivity of 0.23% of the system. In addition, the DOI problem also limits the crystal thickness in this case.
A further prototype scanner, which was published under the name, MADPET, has been developed in its first version in Munich [12]. It is realized with APD's directly coupled to 3.7 mm×3.7 mm×12 mm crystals. This prototype scanner has the problem of increasing the readout channels when using 1-to-1 coupling. In the first scanner, it is not possible to read all channels simultaneously. Moreover, low sensitivity is a problem of the scanner. In a second version of the scanner, MADPET II, this problem was solved, and all APD's can be read out [15]. The second version also has a two-layer readout system, with two layers of crystals with interjacent APD's. Since the crystals are thus divided, DOI positions can also be determined. However, twice the amount of photosensor surface area is also required, and the readout channels are thus increased again. Moreover, as a result of approximately double the amount of photosensors, higher costs are incurred.
The possibility of DOI detection with position-sensitive PMT's has been proven in [10, 17].
Research results with detectors consisting of SiPM's and monolithic crystals are published in [16]. In this approach, SiPM's are used in the same way as the original concept for PMT's and APD's published in [8, 9]. In this method, the photosensors are optically coupled to only one side of the monocrystal. However, the linearly-coded sensor must be extended by a resistor network.
The German patent applications, 102016006056.5 and 102016008904.0, of the applicant disclose sensor chips with which the DOI problem can be solved or reduced.